Heart failure is a debilitating disease in which abnormal function of the heart leads to inadequate blood flow to fulfill the needs of the tissues and organs of the body. Typically, the heart loses propulsive power because the cardiac muscle loses capacity to stretch and contract. Often, the ventricles do not adequately fill with blood between heartbeats and the valves regulating blood flow become leaky, allowing regurgitation or back-flow of blood. The impairment of arterial circulation deprives vital organs of oxygen and nutrients. Fatigue, weakness and the inability to carry out daily tasks may result. Not all heart failure patients suffer debilitating symptoms immediately. Some may live actively for years. Yet, with few exceptions, the disease is relentlessly progressive. As heart failure progresses, it tends to become increasingly difficult to manage. Even the compensatory responses it triggers in the body may themselves eventually complicate the clinical prognosis. For example, when the heart attempts to compensate for reduced cardiac output, it adds cardiac muscle causing the ventricles to grow in volume in an attempt to pump more blood with each heartbeat, i.e. to increase the stroke volume. This places a still higher demand on the heart's oxygen supply. If the oxygen supply falls short of the growing demand, as it often does, further injury to the heart may result, typically in the form of myocardial ischemia or myocardial infarction. The additional muscle mass may also stiffen the heart walls to hamper rather than assist in providing cardiac output. Often, electrical and mechanical dyssynchronies develop within the heart such that the various chambers of the heart no longer beat in a synchronized manner, degrading overall cardiac function. A particularly severe form of heart failure is congestive heart failure (CHF) wherein the weak pumping of the heart or compromised filling leads to build-up of fluids in the lungs and other organs and tissues.
Many patients susceptible to CHF, particularly the elderly, have pacemakers, ICDs or other implantable medical devices implanted therein, or are candidates for such devices. Accordingly, it is desirable to provide techniques for detecting and tracking CHF using such devices. One particularly effective parameter for detecting and tracking CHF is cardiac pressure, particularly LAP, i.e. the blood pressure within the left atrium of the patient. Reliable detection of LAP would not only permit the implanted device to track CHF for diagnostic purposes but to also control therapies applied to address CHF such as cardiac resynchronization therapy (CRT). CRT seeks to normalize asynchronous cardiac electrical activation and the resultant asynchronous contractions by delivering synchronized pacing stimulus to the ventricles using pacemakers or ICDs equipped with biventricular pacing capability. The pacing stimulus is typically synchronized so as to help to improve overall cardiac function. This may have the additional beneficial effect of reducing the susceptibility to life-threatening tachyarrhythmias. CRT and related therapies are discussed in, for example, U.S. Pat. No. 6,643,546 to Mathis, et al., entitled “Multi-Electrode Apparatus And Method For Treatment Of Congestive Heart Failure”; U.S. Pat. No. 6,628,988 to Kramer, et al., entitled “Apparatus And Method For Reversal Of Myocardial Remodeling With Electrical Stimulation”; and U.S. Pat. No. 6,512,952 to Stahmann, et al., entitled “Method And Apparatus For Maintaining Synchronized Pacing”. Reliable estimates of LAP derived from impedance signals would also allow the dosing of heart failure medications (such as diuretics) to be properly titrated so as to minimize the number of episodes of acute heart failure decompensation. Another advantage to providing reliable estimates of LAP from impedance signals is that physicians are more familiar with LAP values. Hence, LAP estimates could be provided to the physician via diagnostic displays, rather than raw impedance signal values, which the physicians might find difficult to interpret.
However, LAP is a difficult parameter to detect since it is not clinically appealing to place a blood pressure sensor directly in the left atrium due to the chronic risk of thromboembolic events, as well as risks associated with the trans-septal implant procedure itself. Accordingly, various techniques have been developed for estimating LAP based on other parameters that can be more safely sensed by a pacemaker or ICD. In this regard, some particularly promising techniques have been developed that use electrical impedance signals to estimate LAP. For example, impedance signals can be sensed along a sensing vector passing through the left atrium, such as between an electrode mounted on a left ventricular (LV) lead and another electrode mounted on a right atrial (RA) lead. The sensed impedance is affected by the blood volume inside the left atrium, which is in turn reflected by the pressure in the left atrium. Accordingly, there is a correlation between the sensed impedance and LAP, which can be exploited to estimate LAP and thereby also track CHF. Another example may be the impedance signals sensed along a sensing vector passing through the lung, such as between an electrode mounted on the a left ventricular (LV) pacing lead and another electrode representing the device case (Case) containing the pulse generator within a subcutaneous thoracic pocket. The sensed impedance is affected by the fluid volume within the lung/thorax, which is in turn reflected and proportional to the pressure within the pulmonary veins that is equivalent to the LAP. See, for example, techniques described in the related patent applications, cited above. See, also, U.S. patent application Ser. No. 11/558,194, by Panescu et al., entitled “Closed-Loop Adaptive Adjustment of Pacing Therapy based on Cardiogenic Impedance Signals Detected by an Implantable Medical Device”, which is incorporated by reference herein.
Although electrical impedance can be used to estimate LAP, it is difficult to reliably calibrate such impedance-based estimation techniques. That is, it can be difficult to accurately and reliably convert detected electrical impedance values into actual LAP values. Accordingly, certain aspects of the above-cited patent applications were directed to providing improved techniques for calibrating impedance-based LAP estimation techniques.
In one example, set forth in the parent patent application (Ser. No. 11/559,235), a linear correlation between LAP and an electrical signal measured within the thorax of the patient is exploited to estimate cardiac pressure. The electrical signal can be, e.g., impedance (Z), admittance (Y), or conductance (G), as measured along a sensing vector passing through the heart of the patient. Note that these electrical parameters are related. Admittance and impedance represent forms of immittance, with admittance being the numerical reciprocal of impedance. Conductance is the numerical reciprocal of resistance. In general, impedance and admittance are vector quantities, which may be represented by complex numbers (having real and imaginary components.) Unless otherwise noted, only the real portion of the impedance or admittance vector is exploited within the equations provided herein. The real component of impedance is resistance. The real component of admittance is conductance. Hence, when exploiting only the real components of these values, conductance can be regarded as the reciprocal of impedance. Likewise, when exploiting only the real components, admittance can be regarded as the reciprocal of resistance.
Suitable conversion factors (also referred to as calibration coefficients) are determined via linear regression, which relate the particular measured electrical signal to LAP, such that subsequent signal measurements can be used to estimate LAP. In one particular example, the conversion factors are “slope” and “baseline” values representative of the linear correlation. Slope may also be referred to as “gain”. Baseline may also be referred to as “offset” or bLAP (i.e. baseline LAP.)
The initial determination of the appropriate slope and baseline conversion factors for use within the patient is referred to as calibration. The conversion factors are preferably re-calibrated as needed to ensure reliable LAP estimates despite anatomical or physiological changes within the patient. That is, the slope and baseline values are recalculated or adjusted, either periodically or on-demand.
The parent application introduced various calibration and re-calibration techniques for use in determining the slope and baseline values. In one illustrative example where the electrical parameter to be measured within the patient is conductance (G), the appropriate slope and baseline values (SlopeG and bLAPG) are determined during an initial calibration procedure based on the assumption that there is a linear relationship between conductance and LAP. To calibrate the slope and baseline values for a particular patient, a “two-point” calibration procedure is employed wherein a first conductance calibration value (G1) and a corresponding first cardiac pressure calibration value (LAP1) are measured within the patient at a first point in time. Then, a second conductance calibration value (G2) and a corresponding second cardiac pressure calibration value (LAP2) are measured at a second point time within the patient. The first and second pressure calibration values (LAP1, LAP2) may be measured within the patient using, e.g., a Swan-Ganz catheter equipped to measure pulmonary capillary wedge pressure (PCWP). The times are chosen such that the first and second cardiac pressure values (LAP1, LAP2) differ substantially from one another (and so the conductance calibration values also differ substantially from one another). In one particular example, the first calibration values (G1, LAP1) are detected while the patient is at rest; whereas the second calibration values (G2, LAP2) are detected while the patient is subject to a condition significantly affecting cardiac pressure, such as isometric muscle contraction, vasodilatation, vasoconstriction, rapid pacing or performance of the Valsalva maneuver by the patient. The slope value is then calibrated by calculating:SlopeG=(LAP2−LAP1)/(G2−G1).The baseline value is then calibrated by calculating:bLAPG=LAP2−SlopeG*G1.Thereafter, LAP is estimated based on newly-detected conductance values using:eLAP=G*SlopeG+bLAPG where eLAP represents the estimated LAP. Alternately, the term zLAP may be used to denote eLAP, where zLAP represents a LAP estimate derived from the impedance signal.
Similar “two-point” techniques may be exploited for calibrating slope and baseline for use with impedance (Z) values or admittance (Y) values. In general, a pair of calibration values, referred to as C1 and C2, can be determined for use with any electrical parameter to be measured, so long as there is a linear relationship between the measured parameter and LAP. Slope is calculated using:Slope=(LAP2−LAP1)/(C2−C1).Baseline is then calculated using:Baseline=LAP1−Slope*C1.Any of the two-point calibration techniques can be expanded to employ multiple data points (i.e. N data points) by exploiting linear regression or other suitable techniques.
In some implementations, particular components of a raw impedance signal (Z0) are exploited, such as a high-frequency “cardiogenic” impedance signal (ZC) representative of the beating of the heart of the patient, a low-frequency respiratory impedance signal (ZR) representative of the respiration of the patient, or an ultra-low frequency circadian impedance signal representative of daily postural/humeral variations of the patient seen in the raw impedance signal (Z0). Corresponding components of the raw admittance signal (Y0) may likewise be exploited. Suitable values for slope and baseline are calibrated for use with the particular signal components to be used.
In another illustrative example set forth in the parent patent application, a cardiogenic pulse amplitude is derived from a cardiogenic impedance signal (ZC) then exploited to estimate LAP. That is, LAP can be estimated based on the cardiogenic pulse amplitude value using appropriate conversion factors. For example, LAP may be estimated using:eLAP=Cardiogenic_Pulse_Amplitude*SlopeCARDIOGENIC+bLAPCARDIOGENIC where Cardiogenic_Pulse_Amplitude is an amplitude value derived from the impedance signal (Z) and SlopeCARDIOGENIC and bLAPCARDIOGENIC are conversion values derived specifically for use converting cardiogenic pulse amplitude values to LAP values. The SlopeCARDIOGENIC and bLAPCARDIOGENIC conversion factors may be calibrated using similar techniques used to calibrate SlopeG and bLAPG. The pulse amplitude extracted from the cardiogenic impedance signal (ZC) may be restricted to certain portions of the cardiac cycle that may be representative of the venous filling phase within the left atrium (i.e., the portion of the cardiac cycle relative to the cardiac electrogram R-wave corresponding to when the V-wave within the LAP waveform occurs).
In yet another illustrative example set forth in the parent patent application, the parameter derived from the electrical impedance signal (Z) is the circadian pulse amplitude value derived from the circadian component of the raw impedance signal (Z0). The circadian pulse amplitude represents the daily postural-dependent thoracic volume variation in the impedance signal and is preferably calculated once per day. Within well compensated heart failure patients, there is typically a significant daily variation in impedance and so the circadian pulse amplitude may be significant, e.g. 20 ohms or more. Within patients with decompensated heart failure, however, there is typically little or no significant daily variation in impedance and so the circadian pulse amplitude is at or near zero. Hence, progression of heart failure correlates with a decrease in circadian pulse amplitudes. As already noted, there is also a correlation with LAP and heart failure, i.e. LAP increases due to progression of heart failure. Accordingly, there is a correlation between decreasing circadian pulse amplitudes and increasing LAP. That is, LAP can be estimated based on the circadian pulse amplitude value using appropriate conversion factors. For example, LAP may be estimated using:eLAP=Circadian_Pulse_Amplitude*SlopeCIRCADIAN+bLAPCIRCADIAN where in Circadian_Pulse_Amplitude is an individual circadian pulse amplitude value derived from the impedance signal over a twenty-four hour period and wherein SlopeCIRCADIAN and bLAPCIRCADIAN are conversion values derived specifically for use converting circadian pulse amplitude values to LAP values. The SlopeCIRCADIAN and bLAPCIRCADIAN conversion factors may be calibrated using similar techniques used to calibrate SlopeG and bLAPG.
Still other LAP estimation and calibration procedures were set forth in the parent patent application, including techniques exploiting signal morphology fractionation parameters. For the sake of completeness, all of these techniques are also described in detail herein below.
Although the estimation and calibration techniques of the parent application are effective, it is desirable to provide still other estimation or calibration techniques. It is to this end that the techniques of the present invention are primarily directed.